1. Field of the Invention
Embodiments of the present invention relate, in general, to radiation detection and particularly to direct conversion x-ray detectors which have embedded electrodes of various composition and use radiation-induced conductivity found in various solid, dielectric materials.
2. Relevant Background
Radiation detectors are used for detection of incoming radiation, such as x-rays, gamma photons and charged/uncharged particles, in a wide range of different applications. For direct detection of photons of various energies, the incoming photons ionize the material of which the detector is made, releasing energetic electrons through interactions such as the photoelectric effect, pair production and the Compton effect. The emitted electrons also cause additional ionization in proportion to the energy of such electrons, which in turn may be detected by a suitable device.
Typically, in radiographic imaging systems, an x-ray source emits x-rays toward a subject or object, such as a patient or a piece of luggage. The beam, after being attenuated by the subject or object, impinges upon an array of radiation detectors wherein the intensity of the attenuated radiation beam received at the detector array is detected. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element which is thereafter transmitted to a data processing system for analysis, ultimately producing an image.
X-ray detectors typically include a collimator for excluding scattered radiation that might be received at the detector, a scintillator adjacent to the collimator for converting x-rays to light energy and a photodiode for receiving the light energy from an adjacent scintillator and producing electrical signals therefrom. In this type of detector, the x-ray energy absorbed by the scintillating material is converted to visible photons which are then directed into a silicon photodiode. The outputs of these photodiodes are converted into digital data by means of various amplifiers followed by analog-to-digital converters and then transmitted to the data processing system for image reconstruction.
A drawback to this indirect approach to x-ray detection is the fact that it is a two step process to convert x-rays into electrical signals that can be further processed for applications such as computed tomography or digital radiography. Also, detectors using scintillator material suffer from the fact that such materials possess memory effects. Visible light that enters a scintillator based detector promptly decays after the cessation of irradiation by x-rays. However this decay is followed by an afterglow effect that may persist for tens of milliseconds. Another drawback of indirect detection is optical cross-talk between two or more detector elements in close proximity. The scintillator material is typically glued to the photodiode array using an optically transparent adhesive. This adhesive is of a finite thickness, thus allowing light, within a certain angle of incidence with respect to the exit plane of the scintillator exiting a certain distance from the edge of such scintillator, to enter the adjacent detector element. This effect can be minimized by making the adhesive as thin as possible, but the integrity of the bond between scintillator and photodiode degrades with a thinner adhesive. Typically, this optical cross-talk effect is the dominant cross-talk mechanism in indirect x-ray detectors.
The direct conversion of x-rays into electrical signals is well known and often employed for dosage and exposure measurement. X-ray detection of this type has two main advantages over the scintillator-photodiode approach mentioned above. First, there is a much quicker decay in the electrical signal after the cessation of irradiation by x-ray energy. Thus the afterglow effect associated with scintillator material is greatly reduced. Secondly, there is simply no need for scintillator material thereby removing the cost of the scintillating material and the cost of assembling such scintillating material into the detector array.
One method for converting x-rays directly into an electrical current is through the use of ion-chambers. Ion-chambers may be constructed by positioning two parallel flat electrode plates a constant distance apart. The plates are typically enclosed in a chamber constructed of a dielectric material such as Plexiglas. The chamber is sealed and filled with an inert gas such as argon or xenon. X-rays are directed in one end of the chamber such that the x-rays pass through the volume of gas between the two parallel plates. The plates are electrically biased so as to create a substantial electrostatic field between the plates. The ionization of the gas by the x-rays in the presence of a large electric field leads to an electric current proportional to the x-ray energy absorbed by the gas. For a constant x-ray energy, the signal may be said to be proportional to the flux of x-ray photons. One of the most significant drawbacks to gas filled ion-chambers is poor x-ray absorption efficiency. Even using chambers filled with xenon gas at high pressure, the absorption efficiency per unit length through such an ion-chamber is poor compared with the scintillator-photodiode approach. Thus ion-chamber detectors are rarely used as x-ray detectors for any type of imaging.
The ion chamber described above is a specific instance of a detector that relies on the radiation induced conductivity of a material that is electrically insulating in the absence of a radiation field. In the case of the ion chamber, the insulating material is a gas, and the presence of a radiation field in the gas lowers the effective electrical resistivity of the gas such that the application of an external electric field causes a significant electrical current to pass through the gas. Others (such as deGaston, U.S. Pat. No. 4,135,090) have used hydrocarbon liquids as the normally insulating material, producing a radiation detector that has similar absorption properties to soft tissue, but is not sensitive to the energy of the detected x-ray.
Another promising direct conversion method in x-ray detection is the use of compound semiconductors composed of materials that have a significantly higher atomic number than silicon. One such material is cadmium zinc telluride (“CZT”). While CZT detectors hold promise, the quality and expense of grown CZT crystals has so far prevented CZT from being used in mainstream x-ray detection.
Whether indirect or direct conversion is used, it is desirable to have not only a measure of the attenuation of the x-rays through a patient or object being imaged, but also a measure of the energy of the x-rays that are not absorbed by the patient or object. This is desirable for determining the composition of the material in the patient or object. This has been accomplished in several ways: 1) The x-ray source energy may be modulated and detector signals recorded for the various x-ray generator tube energies, 2) some portion of the detector array can be masked with a filter that absorbs lower energy x-rays such that the underlying detector of that portion of the detector array responds only to some higher energy portion of the transmitted x-rays, or 3) the detector can be operated in a mode whereby individual x-ray photon events are counted and the size of the respective current pulses produced by a single x-ray photon being absorbed are quantified.
Each such method has its drawbacks. In the case of modulating the x-ray generator tube (method 1 above), the patient or object must receive a higher dose of irradiation, since the detection is done at two different exposures. In the case of masking a portion of the detector array (method 2 above), x-ray energy is needlessly wasted (i.e. not converted into signal) in a portion of the detector. This shortcoming has been minimized by using an entire detector, itself, as the filter such that simultaneous low and high energy signals are created (e.g. by stacking one detector upon another). In the case of photon counting (method 3 above), one must decrease the size of a detection element such that the number of incident x-ray photons per unit time is small enough that one can count the current pulse produced by an x-ray photon without multiple pulses “piling up”, causing the detector electronics to incorrectly classify both the number and the energy of the x-ray photons.